High magnetic field electromagnets have become important elements in various types of equipment over recent years. One important type of such equipment is medical imaging equipment, such as the type commonly referred to as magnetic resonance imaging (MRI) equipment. This equipment utilizes the mechanism of nuclear magnetic resonance (NMR) to produce an image, and accordingly imaging systems operating according to this mechanism are also commonly referred to as NMR imaging systems.
As is well known in the field of MRI, a high DC magnetic field is generated to polarize the gyromagnetic atomic nuclei of interest (i.e., those atomic nuclei that have nonzero angular momentum, or nonzero magnetic moment) contained within the volume to be imaged in the subject. The magnitude of this DC magnetic field currently ranges from on the order of 0.15 Tesla to 2.0 Tesla; it is contemplated that larger fields, ranging as high as 4.0 to 6.0 Tesla, may be useful in the future, particularly to perform spectroscopy as well as tomography. The volume of the subject to be imaged (i.e., the volume of interest, or "VOI") is that volume which receives the high DC magnetic field, and within which the DC field is substantially uniform.
Imaging is accomplished in the VOI utilizing the mechanism of nuclear magnetic resonance in the gyromagnetic atomic nuclei contained therewithin. As such, in addition to the large field DC magnet, the MRI apparatus includes an oscillator coil to generate an oscillating magnetic field oriented at an angle relative to the DC field, and at a frequency matching the resonant frequency of the atoms of interest in the selected volume; frequencies of interest in modern MRI are in the radio frequency (RF) range. As the gyromagnetic nuclei in the defined volume will have a common resonant frequency different from atoms outside of the volume, modulation of a gradient magnetic field (produced by a gradient coil) allows sequential imaging of small volumes. The images from the small volumes are then used to form a composite image of the larger volume, such as the internal organ or region of interest. To produce the series of images, the MRI apparatus also includes a detecting coil in which a current can be induced by the nuclear magnetic dipoles in the volume being imaged.
In operation, as is well known, the magnetic dipole moments of those atoms in the volume which are both gyromagnetic and also resonant at the frequency of the oscillating field are rotated from their polarized orientation by the resonant RF oscillation by a known angle, for example 90.degree.. The RF excitation is then removed, and the induced current in the detecting coil is measured over time to determine a decay rate, which corresponds to the quantity of the atoms of interest in the volume being imaged. Incremental sequencing of the imaging process through the selected volume by modulations in the gradient field can provide a series of images of the subject that correspond to the composition of the subject. Conventional MRI has been successful in the imaging of soft tissues, such as internal organs and the like, which are transparent to X-rays.
It is well known in the art that the spatial resolution of MRI tomography improves as the strength of the available magnetic field increases. Conventional MRI equipment useful in diagnostic medical imaging requires high DC magnetic fields, such as 5 kgauss or greater.
Due to the large number of ampere-turns necessary to produce such high magnetic fields, conventional MRI systems now generally utilize superconducting wire in their DC coils. While the magnitude of the current carried in these coils is extremely high, the superconducting material and accompanying cryogenic systems required in such magnets are quite expensive, and also add significantly to the size and weight of the magnet in the MRI apparatus. In the extreme case, some conventional MRI magnets are sufficiently heavy (e.g., on the order of twenty tons) as to limit the installation of the MRI apparatus to a basement or ground floor laboratory. Addition of the necessary coils or iron required to shield the fringe magnetic field generated by such magnets further increases the size, weight and manufacturing costs of the MRI equipment.
By way of background, U.S. Pat. No. 4,783,628 (issued Nov. 8, 1988) and U.S. Pat. No. 4,822,772 (issued Apr. 18, 1989), both incorporated herein by this reference and commonly assigned with this application, describe superferric shielded superconducting magnets. These magnets described in these patents utilize passive shielding of ferromagnetic material, such as iron. The construction of the magnets described in these patents provide a highly efficient magnet, considering the magnetic field strength as a function of the current conducted in the superconducting loops, and with a highly uniform field in the magnet bore even at very strong magnetic fields such as on the order of 4 Tesla; the shielding is also very good in this magnet, with the 5 gauss line at 50 to 100 cm from the outer wall of the bore.
Another example of a conventional superconducting magnet, but which relies substantially on active superconducting shielding loops is described in U.S. Pat. No. 4,595,899. The magnet disclosed in this reference has a set of three driving coils surrounded by three shielding coils, with the current through the shielding coils adjusted to exactly cancel the dipole outside of the magnet. External ferromagnetic shielding is also located around the shielding coils to assist in further shielding. Examples of other prior magnets used in MRI are described in U.S. Pat. No. 4,612,505, in which shielding is accomplished by way of magnetic soft iron rods, conducting coils, or both; U.S. Pat. No. 5,012,217, issued Apr. 30, 1992, describes yet another prior superconducting magnet utilizing a combination of active and passive shielding.
While actively shielded magnets greatly reduce the magnet weight relative to superferric shielded magnets, the weight of these magnets is still quite substantial, for example on the order of 20 tons. As a result, when used in medical equipment such as NMR stations, the "footprint" required for installation of the magnet and the weight-bearing capability of the floor of the room are both significant, whether the magnet is constructed of the superferric type, the actively shielded type, or a combination of the two. As a result, from the cost standpoint, it is desirable to reduce the physical size and weight of NMR equipment, to reduce the cost of the NMR laboratory.
As indicated in U.S. Pat. No. 4,595,899, and as is true for other conventional electrically shielded magnets, any ferromagnetic shielding used in the magnet is generally located some distance away from the magnet bore. Such placement is intended to limit the effect of iron on the shape and uniformity of the magnetic field in the bore, because, as is well known in the art, iron or other ferromagnetic material near the bore will non-linearly affect the field within the bore, especially at fields above the threshold of magnetic saturation for iron at about 1.0 to 1.3 Tesla. As a result, the sole effect of the iron in these conventional magnets is to provide fringe field shielding at some distance from the magnet, with minimal effect on the field within the bore intended. In some cases, the ferromagnetic shield is located as far away from the bore as to be within the walls of the room surrounding the magnet (or MRI apparatus containing the magnet). This distancing of the ferromagnetic material from the bore causes significant problems in use of the magnets and equipment, either requiring large "footprints" for the magnet and its shielding, or requiring the specially constructed rooms to house the magnet or NMR equipment, either approach resulting in high cost and poor space utilization.
Examples of other prior magnets used in MRI are described in U.S. Pat. No. 4,612,505, in which shielding is accomplished by way of magnetic soft iron rods, conducting coils, or both. In particular, FIG. 3 of U.S. Pat. No. 4,612,505 discloses the use of a pair of relatively large superconducting shielding coils disposed outside the magnet. In addition, FIG. 4 of this reference illustrates a magnet having a shielding sleeve of magnetic soft iron, and shielding coils disposed outside thereof. The magnets disclosed in this reference have relatively low field strengths, such as on the order of 0.23 to 0.3 Tesla, and somewhat high fringe fields, such as 10 gauss or greater at a distance of three meters from the magnet axis.
U.S. Pat. No. 5,012,217, issued Apr. 30, 1992, describes yet another prior superconducting magnet utilizing a combination of active and passive shielding. This reference discloses the placement of a passive ferromagnetic shield around the main driving solenoid, but within the shielding solenoid (which generates the opposing magnetic field). This construction apparently requires that the large mass of the ferromagnetic shield be placed within the cryostat, substantially increasing the cryogenic load and, accordingly, the cost of maintaining the superconducting coils at superconducting temperatures.
In conventional magnets utilizing electrical shielding, either alone or in combination with ferromagnetic shielding, the cost of superconducting material for the outer coils is on the same order as that for the inner driving coils. The cryogenic load is also quite large for superconducting actively shielded magnets, due to the additional superconductors. In addition, it is believed that it is difficult to achieve uniformity of the magnetic field within the bore of the magnet where shielding is accomplished by cancellation of opposing fields, particularly where the desired magnetic field is 1.5 Tesla or greater.
Additional discussion of the effect of iron on the field within the magnet bore is presented in Siebold, et al., "Performance and Results of a Computer Program for Optimizing Magnets with Iron", IEEE Trans. Magnetics, Vol. 24, No. 1 (IEEE, January 1988), pp. 419-422. As particularly noted in FIG. 3 of this article, the coil system must be designed and adapted relative to the iron yoke in order to provide a uniform field in the bore.
The weight and size of the superferric shielded magnets described in U.S. Pat. Nos. 4,783,628 and 4,822,772 can be quite substantial, however, such as on the order of 35 to 130 tons (as compared with actively shielded magnets weighing on the order of 20 tons). As a result, when used in medical equipment such as NMR stations, the "footprint" required for installation of such a magnet, as well as the weight-bearing capability of the floor of the room, are both significant. It is, of course, desirable to reduce the physical size and weight of NMR equipment, thus reducing the cost of the NMR laboratory. Besides the large footprint of conventional NMR magnets, it has been observed that many patients are uncomfortable when placed in magnets of such length, as the patient's entire body is generally disposed within the magnet during much of the imaging procedure. Indeed, conventional cylindrical NMR magnets have been referred to as "tunnel" magnets, representing the sensation perceived by the human subject when placed inside for an imaging procedure. It is therefore also desirable to provide a high magnetic field magnet for purposes of NMR which has good field homogeneity, but where the axial length of the bore is as short as possible.
By way of further background, it should be noted that the driving coils for magnets such as described in the above-referenced U.S. Pat. Nos. 4,783,628 and 4,822,772 are cylindrical in shape, so that a uniform magnetic field is provided over a portion of the axial length of the bore. As described, for example, in U.S. Pat. Nos. 4,587,490 and 4,590,428, and in Everett, et al., "Spherical coils for uniform magnetic fields," J. Sci. Instrum., Vol. 43 (1966), pp. 470-74, it is also known to provide spherical or quasi-spherical coil arrangements to produce a homogeneous magnetic field within the bore.
In addition, it is also known to provide error, or trim, coils in conventional iron-shielded magnets to provide adjustment of the homogeneity of the magnetic field within the bore. One example of such a magnet is described in U.S. Pat. No. 4,490,675, in which the error coils are disclosed as being within the soft iron cylindrical shield. U.S. Pat. Nos. 4,590,428 and 4,587,490 also disclose NMR or MRI magnets including main and error coils within an iron cylinder.
By way of further background, U.S. Pat. No. 4,924,185 discloses another cylindrical superconducting magnet. As disclosed therein, the sense of oppression on the part of the patient is reduced as the ratio of bore length to bore diameter is below 1.90.
By way of further background, U.S. Pat. No. 4,689,591 discloses a superconducting magnet having a plurality of coaxial coils arranged asymmetrically along the axis, resulting in a volume of interest that is offset from the midplane of the magnet. The volume of interest in this magnet, while offset, remains deeply within a cylindrical bore, however, requiring whole body insertion of the patient for MRI procedures.
Another known type of conventional MRI magnet is of the Helmholtz coil type, including such a magnet which utilizes thermally insulated niobium/tin superconducting material. However, this magnet also requires the patient's whole body to be inserted between the Helmholtz coils.
By way of further background, U.S. Pat. No. 5,049,848, issued Sep. 17, 1991, discloses a magnet configuration suited for MRI in mammography. This magnet configuration is of rectangular shape, and includes permanent magnets for generating magnetic flux in two planes in the gap g within which the imaging is to take place. A shimming electromagnet is disclosed as being placed behind the patient, for reducing front edge fringe field.
By way of further background, notched cylindrical coil systems for providing strong magnetic fields are known, as described in M. W. Garrett, "Thick Cylindrical Coil Systems for Strong Magnetic Fields with Field or Gradient Homogeneities of the 6th to 20th Order", J. Appl. Phys., Vol. 38, No 6 (1967), pp. 2563-86. As described in this reference at pages 2578-2583, expansion of ideal magnet elements to larger cross sections by modifying the geometry in an iterative fashion according to Lyle's Principle can be used to arrive at a magnet having a negative current polarity notch or cavity coil within the otherwise cylindrical positive coil; the notch may be at either the inner or the outer radial surface (see FIG. 2 of the reference), or even wholly within the positive coil.
In the method described in the above-cited Garrett article and also according to other conventional methods for designing cylindrical magnets, the designer relies on the property that the axial component of the magnetic field in the bore is a harmonic function within the volume of interest (VOI), which can be expanded into a series of spherical harmonics. The coefficients of this expansion may be expressed as axial derivatives of the axial magnetic field at the origin (center of the VOI). If one assumes that the current density in the cross-section of each magnet coil is constant, these axial derivatives may be calculated directly from coil geometry, without requiring integration of the Biot-Savart Law, allowing the geometry of the magnet to be adjusted so that the undesirable harmonics of the axial field in the VOI vanish. According to this generalized technique, computer-ready methods have been developed with tabular design conditions (see the Garrett article cited above), facilitating the design of such magnets. This general method constrains the location of the VOI to a high degree, however, in order for the calculations to be readily performed; as such, this general method is practically applicable for a VOI centered at the midplane of the cylindrical magnet.
By way of further background, a widely-used conventional method of designing magnets will now be described relative to FIG. 21. FIG. 21 is a cross-sectional representation of a magnet to be designed for generating a magnetic field of desired magnitude, direction and uniformity in a volume of interest VOI. The representation of FIG. 21 is a quadrant of the z-.rho. plane, for a conventional magnet that is both axially symmetric about the .rho.=0 axis, and midplane-symmetric about the z=0 axis. According to conventional techniques, the magnet designer selects the number of coils in the magnet and the shape of the cross-section of each coil. For example, in the magnet of FIG. 21, three coils 70A, 70B, 70C of rectangular cross-section (in the z-.rho. plane) are postulated, along with their initial locations and their current magnitude and polarities.
In the conventional design methodology, the magnetic field in volume of interest VOI is determined, generally by use of a conventional simulation program such as the OPUS computer program, available from Ferrari Associates, Inc. of Jacksonville, Fla. Optimization of the magnetic field parameters toward the design goal is then performed by iterative adjustment of the size, location and current in the coils 70A through 70C, followed by evaluation of the field after each incremental adjustment. Upon the simulated field reaching the optimal condition relative to volume of interest VOI, the position, current and size of each coil is then determined. For the magnet in FIG. 21, for example, the best field (or minimized error) condition is for coils 70A', 70B', 70C' in the position shown.
Comparison of the optimized coils 70A', 70B', 70C' to the initial coils 70A, 70B, 70C in FIG. 21 illustrates the design freedom provided according to conventional magnet design techniques. The number of coils (e.g., three per quadrant) and their shape (e.g., rectangular) remain fixed in this conventional methodology. As such, the magnet design may only adjust the size, aspect ratio, current and location of the coils in obtaining the desired magnetic field. While this number of degrees of freedom is often adequate in the design of cylindrical magnets of modest field magnitude and uniformity requirements where the volume of interest is centered within the bore, this conventional methodology has been observed to be inadequate to successfully design magnets where the volume of interest is not centered within the bore, particularly where the field strength and uniformity requirements are quite stringent.
It is therefore an object of the present invention to provide an electromagnet constructed according to a method having a larger number of degrees of freedom in its optimization.
It is a further object of the present invention to provide such an electromagnet in which the number of coils and their cross-sectional shape (i.e., both the coil boundaries and the number of boundaries) are design parameters and may be selected in optimizing the field strength and uniformity in a selected volume of interest.
It is a further object of the present invention to provide such a method of fabricating a magnet which is applicable to magnets of various symmetry, including cylindrical and planar magnets.
It is a further object of the present invention to provide such a magnet which allows for optimization of the design for a volume of interest at an arbitrary position relative to the bore, including offset from the center of the bore so as to protrude from the bore.
It is a further object of the present invention to provide such a method which allows the center of the volume of interest to be located outside of the magnet bore.
It is a further object of the present invention to provide a superconducting magnet for use in NMR equipment which does not require insertion of the whole body of the patient into the magnet bore.
It is a further object of the present invention to provide such a magnet which is of sufficient field strength to enable in vivo NMR tomography of the internal organs of humans.
It is a further object of the present invention to provide such a magnet which is suitable for the image of specific organs, such as the brain, the female breast, and the like.
It is a further object of the present invention to provide such a magnet which makes the NMR tomography equipment substantially portable.
It is a further object of the present invention to provide an extremely compact superconducting magnet having a high degree of effective shielding.
It is a further object of the present invention to provide such a magnet having a large aperture into which a patient may be placed.
It is a further object of the present invention to provide such a magnet having relatively light weight and low cost.
It is a further object of the present invention to provide such a magnet which can be fabricated using a single cryostat.
Other objects and advantages of the present invention will be apparent to those of ordinary skill in the art having reference to the following specification together with the drawings.